Biosensors

Abstract

A biosensor comprising: first and second substrates (1, 2, 10) defining a cavity therebetween; a sensing structure having a functionalized active surface (4) provided on the first substrate (1, 2) within the cavity; and a flow control structure (11) provided on the second substrate (10) and extending into the cavity, the gap between a distal end of the flow control structure and the sensing structure providing a fluid flow channel across the functionalized active surface (4). and an inlet port (15a) at one end of the fluid flow channel adjacent a proximal end of the flow control structure (11), and an outlet port (15b) at an opposing end of the fluid flow channel, the flow control structure (11) being shaped and configured such that, during use, fluid injected at the inlet port (15a) flows into the fluid flow channel and through the fluid flow channel to the outlet port (15b), thereby exerting a shear force on the functionalized active surface (4).

Description

[Technical Field]

[0001] The present invention relates generally to biosensors, such as graphene-based biosensors, having an active sensor surface, methods for detecting biomolecules using biosensors, and methods for fabricating biosensors. In particular, although not necessarily exclusively, the present invention relates to biosensors having patterned and chemically functionalized graphene surfaces. [Background technology]

[0002] Sensors for detecting biomolecules, called biosensors, are widely known and are used in diagnostic tests such as point-of-care testing (POCT). Point-of-care testing is essentially a diagnostic test performed at or near the patient, the results of which can change the patient's care. Such diagnostic tests are used to detect biomolecules in biological samples collected from patients. A biomolecule, within the meaning of this disclosure, is an organic molecule produced by or occurring in a living organism. The term "biomolecule" includes, but is not limited to, naturally occurring polymeric molecules such as proteins, polysaccharides, and nucleic acids (including synthetic nucleic acids), as well as their analogs, and small molecules such as primary metabolites, secondary metabolites, and natural products.

[0003] In addition to optical and other approaches, many biosensors rely on the general principle of generating an electrical signal when the presence or absence of a biomolecule is detected. Structured semiconductor materials are sometimes used to form channels or other structures on the micrometer scale (microscale) or nanometer scale (nanoscale).

[0004] More recently, graphene has attracted increasing interest for use in biosensors due to its unique physical and chemical properties: its electrical conductivity is 1000 Siemens / meter and its thermal conductivity is 1500-2500 Wm -1 K -1Furthermore, graphene exhibits a wide electrochemical window and low charge transfer resistance, and can be functionalized by adding bioreceptors that affect its reactivity.

[0005] GB2471672B describes a graphene biosensor comprising a patterned graphene layer grown on a SiC substrate. The patterned graphene structure includes at least one channel, with electrical contacts on both sides of the channel, allowing current to pass through the channel. The channel is functionalized with a linker in the form of a receptor that has binding affinity for a target (bio)molecule attached to the graphene surface. One example of such a linker is obtained by attaching nitrobenzene to the graphene surface and then electrochemically reducing it to aniline, but others will be known to those skilled in the art of chemical functionalization / immobilization. In use, one or more target biomolecules are attached to the functionalized channel, and when a current is passed along the channel, changes in the electrical properties of the sensor (induced by the target biomolecules) can be measured.

[0006] This configuration of graphene biosensor is known as a graphene field-effect transistor (GFET). Because graphene has a high surface-to-volume ratio and even minimal concentrations of attached biomolecules alter the channel's conductivity, GFET biosensors are an attractive platform for sensing a wide variety of species, including enzymes, hydrogen peroxide, dopamine, and reduced β-nicotinamide adenine dinucleotide (NADH) molecules. Chemiresistive biosensors are also known, in which the resistance measured by the biosensor increases with increasing concentrations of target biomolecules. Yet another type of graphene biosensor measures the drain-source current and so-called Dirac point (charge neutrality) shift of a liquid-gated GFET upon binding of an analyte (or target biomolecule) to a functionalized graphene surface.

[0007] The main advantages of graphene-based biosensors are their versatility and their sensitivity for detecting a wide variety of different biomolecules or analytes, depending on the functionalization of the graphene surface and the electrical measurements used to detect and quantify them. However, the performance of graphene-based sensors can be adversely affected by several typical problems.

[0008] A significant known problem in this field is known as nonspecific adsorption (NSA). NSA (also known as nonspecific binding or "biofouling") results from the irreversible adsorption of nonspecific biological species (i.e., species other than the target biomolecule of interest) on the active sensor surface, adversely affecting the sensitivity and accuracy of the biosensor and being particularly problematic in POCT instruments that do not support rigorous cleaning of the sensor surface immediately after the species binding step and before electronic detection / measurement of the results. The presence of NSA on the sensor surface can "block" the active area, thereby reducing the sensitivity and selectivity of the biosensor (e.g., giving false-positive results). In some cases, the presence of NSA also acts to increase the baseline, known as "background noise," which in turn reduces the detection limit of the biosensor in that lower concentrations of the target biomolecule may be masked. This noise signal, indistinguishable from the target biomolecule, and any blocking of the active surface area reduces the sensor's performance (i.e., sensitivity, reproducibility, and dynamic detection range). In the laboratory, NSA can be removed from the sensor surface by rigorously washing the sensor with tween-20 phosphate buffered saline (PBS) or PBS alone, but this poses a significant challenge for small, closed POCT systems that cannot incorporate such a sensor washing setup.

[0009] Therefore, there is a continuing desire in the biosensor field to inhibit NSA and / or remove NSA at the sensor surface in order to improve biosensor performance.

[0010] To this end, many methods for NSA inhibition and / or removal have been proposed. Such methods can be grouped into two general categories: passive and active. Passive NSA inhibition or prevention methods can be further classified as physical or chemical passivation. Physical passivation methods are designed to prevent NSA by coating or blocking unreacted active surfaces with blocking agents, such as protein coatings (e.g., bovine serum albumin (BSA)), which inhibit NSA. However, they have been shown to exhibit high lot-to-lot variability, cross-reactivity, and alter the original surface properties (thereby affecting sensor efficacy). On the other hand, chemical passivation methods present laborious functionalization processes, high background signals, and potential damage to the active sensor surface, as well as challenges in maintaining the long-term chemical stability of the active sensor surface. Overall, NSA inhibition using passivation methods, whether physical or chemical, often involves the use of harsh chemicals that are typically unsuitable for many biological applications.

[0011] In contrast, active NSA inhibition / removal methods have emerged as a more promising solution to this problem in the field of biosensors. Such active methods can be classified as transducer-based or fluid-based. Transducer-based methods use electromechanical or acoustic waves for NSA removal. Fluid-based methods, which have not been widely utilized or documented to date, purport to use pressure-driven microfluidic flow to generate shear forces to remove NSA molecules from the sensor surface. The main drawback of transducer-based methods is the need for additional, controlled equipment. On the other hand, fluid-based methods of NSA removal require precise fluid manipulation in a region of interest (ROI) to efficiently remove NSA. Furthermore, some such methods documented in this regard require electrophoretic alignment of non-target species to achieve the desired effect.

[0012] Current biosensor systems rely on either passive or active methods of NSA inhibition and / or removal. However, prior art biosensors have yet to provide an NSA inhibition / removal method that provides an efficient, reproducible, and robust solution to address the above problems.

[0013] Thus, according to a first aspect of the present invention, there is provided a biosensor comprising: first and second substrates defining a cavity therebetween; a sensing structure within the cavity, the sensing structure having a functionalized active surface disposed on a first substrate; a flow control structure disposed on the second substrate and extending into the cavity, a gap between a distal end of the flow control structure and a sensing structure providing a fluid flow channel extending across the functionalized active surface; an inlet port adjacent a proximal end of the flow control structure at one end of the fluid flow channel and an outlet port at an opposite end of the fluid flow channel; A biosensor is provided in which the flow control structure is shaped and configured such that, during use, fluid dropped / injected at the liquid inlet port flows into the fluid flow channel and through the fluid flow channel to the outlet port, thereby exerting a shear force on the functionalized active surface.

[0014] Various aspects of the invention are set out in the independent claims, and additional and / or optional features are set out in the dependent claims appended hereto. These and other features of the invention will become apparent from the following detailed description. [Brief explanation of the drawings]

[0015] Embodiments of the present invention will now be described, by way of example only, with reference to the accompanying drawings, in which: [Figure 1] 1 is a schematic cross-sectional view of a single biosensor according to an exemplary embodiment of the present invention. [Figure 2] FIG. 2 is a schematic cross-sectional view of the bottom package of the biosensor of FIG. 1. [Figures 2A-2C] 1A-1C are schematic plan views of a graphene well of a biosensor according to an exemplary embodiment of the present invention at three respective steps of the fabrication process, namely, with bottom metal contacts (16, 17), top metal clamp (18), and passivation layer (19), respectively. [Figure 3] 2 is a schematic cross-sectional view of the top package of the biosensor of FIG. 1. FIG. [Figure 4] 1 is a schematic cross-sectional view of a multi-biosensor device according to an exemplary embodiment of the present invention. [Figure 5] FIG. 1 is a schematic perspective view of a quarter biosensor device according to an exemplary embodiment of the present invention. [Figure 6] 1 is a schematic cross-sectional view of a multi-biosensor device according to an exemplary embodiment of the present invention. DETAILED DESCRIPTION OF THE INVENTION

[0016] In the following detailed description, embodiments of the present invention are described along with a complete wafer-scale fabrication process for an exemplary graphene-based sensor chipset incorporating novel microfluidics and chipset packaging aimed at providing one or more of the following: i) high sensor specificity while combining the advantages of both passive and active methods of NSA removal on a single chipset, resulting in substantially complete NSA removal; ii) electrostatically assisted binding / immobilization to improve turnaround time (TAT); and iii) real-time multiplexed electrical quantification (MEQ), both chemiresistive and GFET. However, it should be understood that some embodiments may include only one or some of these features, and that while preferred embodiments are described in detail herein, the present invention is not necessarily intended to be limited with respect to any particular combination of described features, except as the scope of the present invention is clearly defined by the appended claims.

[0017] Early detection of disease biomarkers using sensitive, selective, rapid, and cost-effective POCT systems is essential for disease prognosis / diagnosis and real-time patient health monitoring. Attaching specific disease biomarkers to biosensor surfaces is crucial for developing such simple, rapid, and multiplexed sensing platforms for detecting various analytes with high specificity and sensitivity. In recent years, a wide range of immunoassay formats (e.g., ELISA, electrochemical, chemiresistive, optical, magnetic, etc.) have been demonstrated / reported to be useful for the simultaneous detection of multiple analytes from a reference mixture. Furthermore, several methods involving diffusive mixing of analytes using chemical modification of the sensor surface complemented by controlled fluid flow with sophisticated microfluidic channels have been developed to enhance analyte capture efficiency. Despite their performance and capabilities, their incorporation into resource-limited settings (e.g., POCT) requiring simple on-site electronic diagnostic systems is limited by the need for sophisticated electronic / magnetic detection procedures and operational control systems. Recently, carbon nanotubes (single-walled carbon nanotubes, (SW-CNT) or multi-walled carbon nanotubes (MW-CNT)) have shown great potential for use as biosensors. Similarly, graphene, a single-atom thick film of hexagonally bonded carbon atoms, presents potential applications in the field of biosensors by functionalizing its surface with bioprobes for specific biomarkers. Over the past decade, highly sensitive (electrochemical, chemiresistive, and field-effect transistor-based) graphene-based biosensors have been reported.

[0018] Embodiments of the present invention aim to provide a graphene-based biosensor chipset that can address one or more of three key issues: 1) high specificity with (near) absolute removal of NSAs, 2) reduced TAT, and 3) accurate and robust analyte quantification, which have so far hindered the development of robust electronic biosensors for POCT systems.

[0019] Thus, a first aspect of the present invention addresses NSA inhibition / removal, aiming to significantly reduce or substantially eliminate NSA molecules from the ROI, i.e., the sensor surface. In exemplary embodiments of the present invention described below, this is achieved through unique open / closed loop microfluidic channels and the use of gentle electrical (ac) or electromagnetic stirring, where the microfluidic channels generate enhanced hydrodynamic shear forces and directional flow of fluid across the region of interest. As known to those skilled in the art, blocking agents can also be used to further enhance NSA inhibition.

[0020] A second aspect of the present invention aims to improve turnaround time (TAT). It may be apparent that minimizing TAT in biosensors, particularly POCT devices, is an important factor, and this is achieved in the following exemplary embodiment by using electrostatic coupling and improving the speed of such coupling by applying an electrostatic (DC) potential to capacitive electrodes embedded below the functionalized sensor surface.

[0021] A third aspect of the present invention addresses the need for (near) real-time MEQ and uses both chemoresistance and GFET measurements for this purpose. (Near) real-time MEQ can be obtained utilizing a novel processing module configured to collect chemoresistance and Dirac point shift data upon analyte binding and to use signal processing techniques for analytical quantification of target bioanalytes.

[0022] In the following detailed description, any reference to directional terms such as "top," "top," "bottom," "lower," "below," "side," etc., is understood to be used (and applied) solely in relation to the orientation of the device as shown in the accompanying drawings, and these terms are in no way intended to be limiting with respect to the orientation of the device in use.

[0023] Referring to FIG. 1 of the drawings, a schematic cross-sectional view of a biosensor device according to an exemplary embodiment of the present invention is provided. The example device comprises a graphene sensor that can be used for multiplexed electrical quantification (MEQ) of bioanalytes, with provisions for electrically / electromagnetically assisted fluid agitation and forced fluid propagation across a region of interest within an open / closed microfluidic channel, along with electrostatic biasing of the graphene for improved turnaround time (TAT). For wafer-scale fabrication purposes, the device comprises two separate components: a bottom package 100 and a top package 200. As described in more detail below, the bottom package 100 includes the graphene FETs, microfluidic channels, inlet reservoirs, wells / cavities, and outer reservoirs, while the top package 200 includes a solid substrate (e.g., glass or other optically transparent / semitransparent material), a trapezoidal dome, and polymeric microfluidic channels, fluid control valves, and a sample delivery system. When integrated with the "bottom" package 100, the "top" package 200 forms a complete sensor platform for MEQ of bioanalytes. In certain exemplary embodiments described herein, reference is made to graphene-based biosensors for detecting cardiac biomarker proteins, although it will be understood that the invention is not intended to be limited in this respect.

[0024] Further, with reference to Figures 2 and 2A-2C of the drawings, the device's "bottom" package 100 comprises a graphene FET-based biosensor. In the method of fabrication of the "bottom" package 100, CVD-grown graphene 4 is transferred directly onto pre-deposited metal contacts (including an Ag / AgCl reference electrode 12), while depositing graphene on a commonly found SiO2 / Si substrate, or (in this particular exemplary embodiment) on a self-assembled monolayer (SAM)-coated dielectric / insulating SiO2 / Si substrate 1, 2 (for the purpose of improving sensor system performance, such as lowering the graphene sensor's Dirac point). As can be seen more clearly in Figure 2A of the drawings, a graphene channel 4a is patterned after the graphene transfer and cleaning process. As known to those skilled in the art of microdevice fabrication, graphene FET-based biosensors can be reliably fabricated, for example, in a research-oriented cleanroom, and the CVD graphene channel 4a effectively functions as the active channel for bioanalyte detection. As shown in Figure 2A of the drawings, an electrode 16 is located on the substrate 2 "below" each graphene channel 4a at one end, and a metal contact 17 extends "below" all of the graphene channels 4a at the opposite end. In Figure 2B, a top metal clamping layer 18 can be seen, clamping each end of the graphene channels 4a to a respective electrode 16 at one end and to a contact 17 at the other end. In Figure 2C, a passivation layer 19 can be seen, covering all of the structures on the substrate 2 except for the graphene channels 4a and the reference electrode 12.

[0025] The graphene channel (generally designated "4" in Figures 1 and 2 of the drawings) is fixed with a suitable linker 5, and the probe molecule 6 used will depend on the analyte to be detected (e.g., in this case, a cardiac biomarker protein). Suitable methods for chemically functionalizing the graphene channel are described, for example, in GB Patent No. 2471672 and will be known to, inter alia, those skilled in the art. Therefore, this aspect of the fabrication method will not be discussed in further detail herein.

[0026] Multiple such "all-graphene" FETs can be fabricated on a single substrate and fluidically coupled together by microfluidic channels 15 so that analytes 8 (fluids) can be delivered onto the active (functionalized) graphene regions of each device, as shown and described in more detail below with reference to Figure 4 of the drawings. In Figure 1 of the drawings, microfluidic channels 15a define the inlet ports for the illustrated biosensors, defined between a first photoresist structure 9a (e.g., any biocompatible polymer, such as SU8) on substrate 2 and substrate 10 adjacent the proximal end (base of largest diameter) of dome 11 on top package 200 (described below with reference to Figure 3 of the drawings). A second polymer structure 9b is provided on the opposite edge of functionalized active surface 4 and defines an outlet port 15b between its distal end and substrate 10 of top package 200. Thus, during use, fluid enters the biosensor through inlet port 15a and flows through a fluid flow channel defined between the distal end (smaller diameter base) of dome 11 and the functionalized active surface 4 to outlet port 15b. A third photoresist / polymer structure 9c is provided on substrate 2 spaced from second photoresist / polymer structure 9b, such that an outer reservoir 29 is defined between the second photoresist / polymer structure 9c and the third photoresist / polymer structure 9b. As described in more detail below with reference to Figure 3, a porous polymer track 13 is provided on substrate 10 of top package 200, extending into outer reservoir 29, and a pinhole (approximately 50-100 µm) is provided in glass substrate 10 of top package 200 adjacent to the end of outer reservoir 29 (i.e., near where third photoresist / polymer structure 9c is located and downstream of outlet port 15b).

[0027] With further reference to Figure 3 of the drawings, the "top" package 200 of the biosensor device comprises a solid (beneficially transparent / semi-transparent, e.g., glass) substrate 10, a trapezoidal dome 11, an Ag / AgCl reference electrode 12' on the dome 11, and a small-diameter pinhole 14 (approximately 50-100 μm) in the glass. The customized dome 11 can be formed from a polymer, such as PDMS or SU8, on the glass substrate 10 using, for example, spin-coating / screen-printing techniques that would be well known to those skilled in the art. Additionally, a highly porous polymer may be deposited (e.g., using spin / spray coating or screen-printing) on ​​the edge of the solid substrate 10 to incorporate the polymer into the outer reservoirs 29 (see Figure 4) of the multi-sensor package. This polymer helps ensure unidirectional fluid flow through the device and avoid backflow due to liquid adsorption. Furthermore, the porous polymer layer prevents fluid flooding within the sensor.

[0028] Referring again to FIG. 4 of the drawings, a top package 200 is integrated with the bottom package 100 to provide a multi-sensor mechanism, with the sensors fluidly coupled together by microfluidic channels 15. The top package 200 is aligned on top of the bottom package 100 so that each trapezoidal dome 11 is precisely aligned "above" its respective graphene channel 4a, with a small gap between it and the porous polymer track 13 in the outer reservoir 29. The microfluidic configuration further includes an inlet 27, a well 28 (between the dome 11 and each functionalized sensor surface), and an outer reservoir 29 (downstream of each sensor). A plasma separation membrane 25 is provided across the inlet 27. The microfluidic channels 15, inlet 27, well 28, and outer reservoir 29 can be fabricated using, for example, a biocompatible epoxy-based SU8 photoresist. The wells 28 may advantageously be created by a first SU8 deposition and the interconnected microfluidic channels 15 by a second SU8 deposition, although the invention is not intended to be necessarily limited in this respect. Any suitable polymer may be used for microfluidic fabrication, and the invention is in no way limited to the use of SU8. Other suitable polymers will be apparent to those skilled in the art.

[0029] In this particular exemplary embodiment of the present invention, a biotin-streptaavadin-biotin based sandwich type chemistry is used to immobilize probe molecules 6 onto graphene 4. Certain probes (e.g., for cardiac biomarker detection as in this case) are immobilized onto graphene using such linking chemistries as is well known to those skilled in the art.

[0030] Gentle alternating current electrohydrodynamic or electromagnetic agitation may be applied to each active sensor surface 4 via the electrodes described above in a manner well known to those skilled in the art of electrochemistry. Any NSA molecules formed on the sensor surface will be “loosened” by such agitation. In one exemplary method, NSA removal may be performed separately from the testing process or integrated into the testing process, depending on, among other factors, the type of test being performed and the required accuracy and specificity. Thus, in an exemplary method, a sample fluid is dropped / injected into the device via inlet 27 and flows through microfluidic channel 15. The dropped / injected sample fluid may be mixed with a diluent fluid, such as phosphate-buffered saline (PBS), before being dropped / injected into the sensor device. Alternatively, the sample fluid may be dropped / injected first, followed by the injection of a diluent fluid that mixes with the sample fluid and causes the mixture to flow through the sensor and across the sensor surface. Target molecules in the testing fluid bind to probe molecules 6. Any NSA molecules formed on the sensor surface are "loosened" by the agitation and pulled away from the sensor surface by fluid flow through the device (and across the sensor surface). In other words, pressure-driven fluid flow through the microfluidic channels 15 is used to pull the NSA molecules away from the active sensor surface. The fluid then flows into the respective outer reservoirs 29, with the porous polymer blocks 13 acting to prevent any backflow of fluid in the opposite direction due to liquid adsorption. Of course, a blocking agent (e.g., a protein coating such as bovine serum albumin (BSA)) may also be provided on the active sensor surface to inhibit NSA thereon. Such passive methods of NSA inhibition are well known to those skilled in the art and would be useful in enhancing the NSA inhibition / removal method of this exemplary embodiment of the present invention.

[0031] In the above-described embodiment, the diluted test fluid acts to remove NSA. However, in alternative embodiments, dilution of the sample fluid within the sensor device may occur separately from the NSA removal step. In this case, as described above, after dilution of the test sample, a fluid (such as PBS) may be injected into the device to effect NSA removal. This may be accomplished through the same microfluidic channel 15 used to deliver the test sample and dilution fluid to the active sensor surface. In this case, and in some embodiments, one or more microfluidic valves may be required to prevent backflow and cross-flow mixing between the two fluids. In alternative embodiments, a separate microfluidic channel mechanism may be provided.

[0032] As discussed in the previous section, a description of calculations and simulations for the microfluidic design of certain exemplary embodiments of the present invention follows to demonstrate its practical feasibility. Shear stress and shear force are important parameters for detaching NSA molecules from the active surface walls, and aspects of the present invention offer the potential to exert significant control over these properties using microfluidics. Appropriate shear stress must be applied while ensuring that the flow regime across the active sensor surface is laminar. Laminar flow and shear stress can be achieved and controlled by adjusting the channel height using microfluidic design. Figure 4 shows an exemplary quarter-wave biosensor device integrated with microfluidics, as described above. The channel dimensions are reduced at the location of the sensor surface by the dome 11, and the trapezoidal dome 11 is configured to generate shear force and maintain laminar flow (Re<2300). The height difference between the target area (active surface area) and the dome structure and its dimensions defines the generated shear force and pressure. The smaller the height difference, the greater the shear force generated within the channel. Furthermore, shear force can also be controlled by the inlet flow rate. Theoretical calculations and modeling have shown that changing the channel height affects shear stress more than changing the flow rate. As mentioned above, to generate higher shear stress, the height difference needs to be small, but the flow should also be laminar in nature. The shear force required to remove NSA molecules depends on the bond dissociation energy between the analyte and probe molecules. Typically, due to specificity, the bond dissociation energy is higher between the probe and analyte molecules, and a higher shear force would be required to separate the linker, probe, and analyte than would be required to remove the NSA molecules, especially if the NSA molecules have already been "loosened" by AC electrohydrodynamic or electromagnetic agitation. Therefore, any remaining nonspecific molecules present in the analyte solution will weakly bind (physisorb) to the sensor surface. These physisorbed NSA molecules can be relatively easily detached from the sensor surface with less shear force than would be required to remove the probe and analyte molecules.The shear force required will, of course, depend on the type of linker, probe, and analyte used in fabricating the sensor surface, but the goal may be to generate sufficient shear force to remove the NSA molecules.

[0033] Theoretical models demonstrate that when R / h<0.25, the fluid shear stress on the cell is equal to the shear stress at the wall, where R is the cell diameter and h is the height of the microfluidic channel. The wall shear stress, or simply shear stress, required to remove an NSA molecule depends on the total binding force of the cell. The channel height required to remove a molecule of a certain diameter can be estimated from theoretical calculations. It has been observed that smaller channel heights result in higher shear stresses on the molecule. Computational modeling can be performed to determine the dimensions and shear forces required to remove, for example, NSA molecules. The corner and edge dimensions of the dome can be optimized to minimize pressure drop within the well. Multiphysics-based modeling can be used to create suitable microfluidic designs using domes.

[0034] In an exemplary testing method, sample fluid may be injected (through the inlet to the sensor device) over a period of time, e.g., 15 seconds. Due to diffusive broadening just before the inlet section, a smooth pulse enters the sensor, which can be described by a Gaussian distribution at the flow cell inlet. The design described and illustrated in the drawings results in greater shear stress across a narrow region of interest (i.e., across the sensor surface) and lower shear stress in the remainder of the channel due to differences in channel height. The bond dissociation energy between the linker, probe, and analyte molecules varies depending on their chemical nature. An NSA removal shear stress of approximately 2-10 Pa is required to ensure surface cleaning without dissociating target molecules from the sensor surface. The gap between the dome and the ROI can be precisely adjusted to achieve the shear force necessary to maximize NSA removal while keeping the probe-analyte bond intact and / or dissociating immobilized probes from the sensor surface. As explained above, controlling the shear force can remove all weakly bound NSA molecules from the surface. The required shear force can be estimated from theoretical calculations and depends on the type of linker and probe molecule used for detection, as will be apparent to those skilled in the art. Generally, the shear force required to remove an NSA molecule can be approximately 50% of the adhesion strength. The shear force can be adjusted according to requirements by changing the dimensions of the microfluidic channel, particularly the height difference between the dome and the active channel.

[0035] As mentioned above, the device illustrated and described herein not only functions to remove NSA but also provides directional flow of analyte fluid. Directional flow (preventing backflow) is essential to avoid contamination of the analyte with NSA molecules. Otherwise, the sensor may exhibit a lot of background noise, making it difficult to discern the original analyte signal. The pinhole 14 in the glass substrate 10 of the "top" package 200 helps prevent blockage of analyte fluid flow within the channel or well, which can be caused by internal air pressure buildup. The pinhole 14 can be covered with a layer of gas-permeable polymer (e.g., a thin PDMS polymer) that allows only air to pass through while blocking fluid. As mentioned above, this complete top package 200 must be aligned with the bottom package 100, thereby creating a fluidic channel for analyte fluid flow. The described device design is also suitable for immobilizing linkers, probes, and blocking agents during the manufacturing process, as well as for facilitating the aforementioned NSA removal during normal use.

[0036] Thus, during the manufacturing process, the top and bottom packages 100, 200 are fabricated separately. The graphene channels 4 of the bottom package can be aligned with respective microneedle-based nozzles to deliver a volume of functionalization fluid to the active surfaces within the wells / cavities before completing the packaging process by aligning the top and bottom packages. Using a device of this design, it is possible to simultaneously functionalize each well / cavity without cross-contamination. This makes it possible, for example, to provide multiplexed biosensor devices with multiple wells / cavities, each containing multiple graphene channels and whose graphene surfaces are functionalized to attract different target molecules. This can be achieved at the nanoscale and with a very fast fabrication process.

[0037] One of the key requirements for methods of removing NSA molecules using shear force is maintaining a substantially uniform shear force across the entire region of interest with minimal pressure drop. Ideally, there should be little or no pressure drop within the channel to effectively remove NSA molecules. However, there is always a trade-off between the channel length and the pressure drop in the region of interest (ROI). Longer channel lengths result in significant pressure drops within the channel, which is not conducive to uniform NSA removal. The channel height is reduced in the region of interest, as described above, to achieve higher shear forces while ensuring that the flow remains laminar. This invention combines microfluidic technology with electrical agitation of the analyte fluid to effectively remove NSA molecules. The hydrodynamic shear force can be adjusted by varying the dimensions of the microfluidic channel across the ROI within the biosensor. Simultaneous electrical pulse agitation helps remove NSA molecules from the surface.

[0038] Therefore, in summary, the majority of NSA molecules are physically adsorbed or weakly bound to the sensor surface compared to target analyte molecules. Generally, target analyte molecules bind to specific probes / receptors through strong chemical interactions. The bond dissociation energy is higher between specific probes and analyte molecules, requiring higher shear forces to dissociate the bond compared to physically adsorbed NSA molecules. By adjusting the appropriate shear force, weakly bound NSA molecules can be removed from the surface. In an exemplary embodiment of the present invention, the required shear force is optimized for a biotin-streptavidin-based linker. The described design not only removes NSA but also provides a directional flow of the analyte liquid. This aspect of the present invention provides an efficient method for removing NSA and separating NSA molecules from an active surface.

[0039] In prior art biosensors, immunosensor quantification can be achieved by collecting average independent electrical signals from chemiresistive, electrochemical, and / or field-effect transistor sensors. However, prior art biosensor structures do not provide a unique scheme in which two or more types of signals from a single sensor device are collected and used to provide robust average quantification. In embodiments of the present invention, a device structure provides multiplexed electrical quantification (MEQ) while measuring chemiresistive and liquid top / bottom-gate G-FET signals on a single device platform. Thus, robust average quantification of measured bioassay data (average of chemiresistive + liquid-gate FET signals) is presented. Real-time MEQ can be achieved by a processing module configured to collect multiple data upon analyte binding and perform signal processing to generate an average % change output as analytical quantification of the target bioanalyte.

[0040] Achieving fast sensor response times, or small turnaround times (TAT), is another important challenge, especially in the field of point-of-care (POCT) devices used to diagnose critically ill patients and / or emergency biosensors where speed and accuracy are crucial. The TAT primarily depends on the chemical reaction kinetics between the biological species (probe and analyte molecules) and the graphene surface, which is further defined by the diffusion rate of the biological species in the dilute solvent.

[0041] In embodiments of the present invention, the binding / immobilization of biological species (probe and analyte molecules) can be accelerated by electrical biasing of electrodes embedded under the active graphene channel area. The opposing charges (electrical bias) present on the electrodes help improve reaction kinetics while attracting biomolecules from the solvent solution toward the graphene surface. This electrostatically assisted probe / analyte binding results in a short turnaround time (TAT), which is highly desirable for POCT devices.

[0042] Thus, the graphene biosensor devices described above (and representing exemplary embodiments of the present invention) can provide increased biosensor sensitivity, selectivity, detection limits, and reproducibility, as well as increased background noise suppression, compared to prior art devices.

[0043] It is highly desirable to minimize manual intervention in POCT systems while performing clinical diagnostics. It is the intent of embodiments of the present invention to provide a robust automated system for the actual sample collection, dilution, delivery, and performance of sensitive detection of specific analytes while achieving maximum NSA removal through the proposed active NSA removal scheme described above.

[0044] Referring to Figure 6 of the drawings, a sensor chip package having an alternative top package configuration is shown schematically. It should be understood that where features are common to those of the above-described embodiments, like reference numerals are used in Figure 6. The package shown in Figure 6 includes a pair of microvalves that can be actuated using electromagnetic / electromechanical means, thereby providing a level of automation for sample mixing, dilution, and controlled NSA removal via air / fluid pressure. The benefits of automation include precise control of fluid flow, the pressure required for NSA removal, and TAT.

[0045] The description of the biosensor package in use is as follows: 1. Original Sample Delivery and Dilution: The sample delivery platform may include a first O-ring 33, which in this exemplary embodiment may be made of PTFE or another suitable material, having an inner diameter of about 2-3 mm and an outer diameter of about 5-7 mm, supported (glued) on a second O-ring 35 of soft rubber having an inner diameter of about 2-3 mm and an outer diameter of about 5-7 mm. The first and second O-rings 33, 35 may further support a tapered polymer pipette 36 of a volume appropriate for holding a blood sample of about 2-10 μL. The first O-ring 33, second O-ring 35, and pipette 36 are all part of a substrate assembly. The functions of these components are as follows: a. First O-ring 33 (eg, PTFE / Teflon washer): acts to collect blood from the fingertip and direct it into pipette 36. b. A second O-ring 35 (eg, soft rubber) acts as a cushion for the first O-ring 33 (eg, PTFE / Teflon washer) and (eg, PTFE) pipette 36. c. Pipette (like structure) 36: serves to hold the blood sample and filter the plasma through the plasma membrane 25. The membrane 25 inside the pipette also serves to hold the dilution solvent (e.g., PBS) in the reservoir 21 and prevent leakage through the pipette during transport. The membrane 25 may be replaced with a porous membrane depending on the type of application (e.g., for analyzing wash water in a food processing plant for the detection of allergens in food processing, or for analyzing any other sample not involving blood, where a plasma membrane may not necessarily be desirable). The volume defined by the inner diameter of the first O-ring 33 and the portion of the pipette 36 above the membrane 25 may define the volume of the original sample before dilution.

[0046] A blood sample can be delivered directly by touching a punctured finger at the center of the first O-ring 33. Plasma filtered through the plasma membrane 25 can mix with a pre-stored dilution medium (e.g., PBS) in the dilution channel / reservoir 21 to provide a predetermined sample concentration. Meanwhile, a blood-free sample can be delivered directly into the first O-ring 33 by a micropipette, such that it travels through the (in this case porous) membrane 25 and mixes with the dilution fluid in the reservoir 21. The dimensions and design of the reservoir 21 can determine the volume of the dilution fluid; therefore, the reservoir 21 can be designed to achieve a sample dilution of 100, 200, 500, etc., with a given volume of original sample defined by the volume of the pipette. Alternatively, a fixed volume of dilution fluid can be delivered to the channel / reservoir 21 from a separate reservoir constructed in the top package (not shown in FIG. 6) by an electromagnetically / electromechanically actuated valve (similar to the first and second valves 30, 31 shown in this embodiment and described below). The original sample volume in the pipette can also be accurately estimated by non-contact optical measurement of the volume, as described, for example, in Chinese Patent No. CN104132613. The non-contact optical volume measurement components can be integrated with an electronic readout system. In this case, all critical components of the sample delivery platform need to be biocompatible.

[0047] 2. Delivery of the diluted sample: The first microvalve 30, which can be made of a ferromagnetic material or any polymeric material (e.g., PTFE / Teflon, etc.), is designed to initially block the delivery of the diluted sample to the sensor located in the cavity 28. The first valve 30 is then activated by an electromagnetic / electromechanical switch installed in the electronic readout system to open the valve and control the fluid flow. The electronic activation (opening / closing) of the first valve 30 may be programmed through a software program integrated into the data analysis scheme, so that automation of the fluid flow can be achieved. In practice, a given flow rate of the diluted sample and the temporary switching of the first valve 30 can define the sample volume (μl) delivered to the sensor surface.

[0048] 3. NSA Removal by Controlled Pressure: NSA removal is achieved by shear force across the ROI. The appropriate shear force can be achieved by spraying pressurized air / fluid through a piston-shaped second valve 31. The second valve 31 can be made of ferromagnetic or any polymeric material (e.g., PTFE / Teflon, etc.) and can support a soft polymeric round disk 37. The appropriate amount of fluid can be hidden within an enclosure 38 in the top package, which holds the piston valve 31 at one end and a porous membrane 39 at the other end. The porous membrane 39 is permeable to the pressurized air / fluid. Using an automated manufacturing process on an industrial scale, during the manufacture of the top package, the first O-ring 33, the second O-ring 35, and the membrane 25 can be fitted to the pipette 36. The enclosure 38 can be filled with air / fluid and the reservoir 21 can be filled with a diluent fluid. Again, actuation of the second valve 31 can be performed by an electromagnetic / electromechanical mechanism, as described above in paragraph 2. The purpose of this second valve 31 is to provide the necessary NSA effect while forcing air / fluid into the chamber through the membrane 39, which helps generate the desired pressure in the ROI to remove the NSA from the sensor surface (in this case, graphene). The forced / flushed fluid exiting the channels and cavity in the bottom package can be absorbed by the porous membrane 13 located in the reservoir 29. The advantage of using air for NSA removal instead of excess fluid (or sample fluid) is that it prevents saturation of the porous membrane 13 with excess fluid and avoids potential backflow of fluid from the reservoir 29 into the cavity 28. Again, automatic actuation of the second valve 31 can be achieved using a software module, similar to the first valve 30 for sample delivery (described above in paragraph 2). The programmable electromagnetic / electromechanical actuation of these microvalves 30, 31 can provide automation of sample dilution, dosing, and NSA removal effects, which is highly desirable for small, handheld POCT systems.

[0049] From the foregoing description, it will be apparent to those skilled in the art that modifications and variations can be made to the described embodiments without departing from the scope of the invention as defined by the appended claims.

Claims

1. A biosensor comprising: - first and second substrates defining a cavity therebetween; - within said cavity, a sensing structure having a functionalized active surface provided on said first substrate; a flow control structure disposed on the second substrate and extending into the cavity, the gap between the distal end of the flow control structure and the sensing structure providing a fluid flow channel across the functionalized active surface; an inlet port at one end of the fluid flow channel adjacent the end of the flow control structure opposite the distal end, and an outlet port at the opposite end of the fluid flow channel; a flow control structure shaped and configured such that, in use, fluid dripped / injected at the inlet port flows into the fluid flow channel, through the fluid flow channel to the outlet port, thereby exerting a shear force on the active surface; The biosensor comprises: a first structure on the first substrate located at a first end of the functionalized active surface and extending into the cavity, the first structure comprising a biocompatible polymer, wherein a gap between a distal end of the first structure and the second substrate defines the inlet port; and a second structure on the first substrate downstream of the functionalized active surface, the second structure comprising a biocompatible polymer, wherein a gap between a distal end of the second structure and the second substrate defines the exit port; and a third structure located on the first substrate, spaced apart from and downstream of the second structure and the exit port, the third structure comprising a biocompatible polymer, wherein a space between the second structure and the third structure defines an outer reservoir; and a porous polymer block extending from the second substrate into the outer reservoir; A biosensor, wherein the sensing structure comprises a graphene layer functionalized with linker and probe molecules configured to bind to analyte molecules of interest.

2. 10. The biosensor of claim 1, wherein the flow control structure comprises an outer surface adjacent the inlet port that is rounded and convex relative to a fluid flow path from the inlet port to the outlet port.

3. 3. The biosensor of claim 2, wherein the flow control structure may comprise a trapezoidal or truncated dome having first and second bases that are substantially parallel and planar, the diameter of the first base being larger than the diameter of the second base, the first base being on the second substrate, the second base being located closest to the functionalized active surface, and the gap being between the second base and the functionalized active surface.

4. The biosensor of claim 1 , wherein the second substrate is a solid dielectric substrate.

5. The biosensor of claim 1 , wherein the flow control structure comprises a polymer.

6. 10. The biosensor of claim 1, wherein a small diameter hole is provided in the second substrate and extends into the outer reservoir downstream of the porous polymer block.

7. The biosensor of claim 1 , further comprising an electrode positioned between the sensing structure and the first substrate for applying an electrostatic potential to the sensing structure.

8. 8. The biosensor of claim 7, wherein the electrode is located at one end of the sensing structure and a conductive contact is located at an opposite end and downstream of the sensing structure, the conductive contact being provided between the first substrate and the sensing structure.

9. 8. The biosensor of claim 7, further comprising means for applying an electrostatic potential to the electrodes, the electrostatic potential being of a polarity that acts to attract analyte molecules of interest.

10. 10. The biosensor of claim 1, further comprising a quantification module arranged and configured to, during use, receive an electrical signal from the sensing structure and determine the presence and / or concentration of analyte molecules of interest in a sample fluid flowing through the fluid flow channel.

11. 11. The biosensor of claim 10, wherein the quantification module is configured to determine the presence and / or concentration of analyte molecules bound to the functionalized active surface from the chemiresistivity and GFET signals based on an average amount derived from the chemiresistivity and GFET signals.

12. 12. The biosensor of claim 11, wherein the quantification module is configured to generate an average % change in the combined chemiresistivity and GFET signal expressed as an analytical quantification of the analyte molecule of interest.

13. The biosensor of claim 1 , wherein the second substrate is a solid optically transparent / semi-transparent dielectric.

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